1. Field of the Invention
This invention relates to cages for facilitating the fusion of adjacent bones or adjacent bone surfaces, and more particularly to degradable cages for spinal interbody fusion.
2. Description of the Related Art
Back pain resulting from instability of the spinal system is a rapidly growing condition in the United States. Spinal fusion procedures are expected to grow from over 400,000 procedures in 2004 to 550,000 procedures in 2010. This is driven by an aging population, increasing obesity, and increased patient education and awareness of the fusion procedures. While current segmental spinal fusion relieves pain by eliminating spinal instability, complications associated with conventional metallic cages, including; difficulty of revisions, increased adjacent level disc disease due to increased loading, implant migration or failure, imaging artifacts, stress shielding, and limited bone grafting significantly reduce the efficacy of the interbody fusion. Non-degradable polymeric materials such as polyetheretherketone (PEEK) have been introduced as new cage materials as they are radiotransparent and compliant and can enhance postoperative image modality and fusion rate. However, since clinically reliable reports of using these cages are scarce, concerns still remain that synovitis and the lymphatic spread of non-absorbable polymer debris may be found after intra-articular procedures (see Cho et al., “Preliminary experience using a polyetheretherketone (PEEK) cage in the treatment of cervical disc disease” Neurosurgery 52(3):693 2003 and Neurosurgery 51:1343 2002; and Parsons et al., “Carbon fiber debris within the synovial joint. A time-dependent mechanical and histologic study”, Clinical Orthopaedics & Related Research 1985:69-76).
It has been reported that spine musculoskeletal impairments, including degenerative disc disease, stenosis, spondylolysis, and/or spondylolisthesis, represent more than one-half (51.7% or 15.4 million incidents) of the musculoskeletal impairments reported in the United States. In the United States, 279,000 spinal arthrodesis were performed in 1990, with 26 lumbar fusions performed per 100,000 people (see Andersson, “Epidemiological features of chronic low-back pain”, Lancet 354:581-5, 1999). In 1995, approximately 160,000 spinal fusion surgeries were performed (see Praemer et al. “Musculoskeletal Conditions in the United States” Park Ridge: American Academy of Orthopaedic Surgeons, 1999). A recent report in 2000 (see Sanhu, “Anterior lumbar interbody fusion with osteoinductive growth factors”, Clinical Orthopaedics and Related Research 371:56-60, 2000) revealed that in the United States alone approximately 360,000 patients underwent some type of spinal arthrodesis. The use of cage devices has become an adjunct to interbody fusion for degenerative disorders of the lumbar spine. However, current metallic cages are associated with excessive rigidity that increases incidence of postoperative complications such as stress-shielding, the migration or dislodgement of the cage, pseudoarthrosis, or the combined adverse symptoms (see van Dijk et al. “The effect of cage stiffness on the rate of lumbar interbody fusion: An in vivo model using poly(L-lactic acid) and titanium cages”, Spine 27:682-8, 2002). Metallic cages can also interfere with visual assessment of arthrodesis and the integrity of the spinal canal and neural foramina due to image artifact. The stress-shielded environment resulting from excessive metallic cage stiffness lowers intracage pressure (see Kanayama et al., “In vitro biomechanical investigation of the stability and stress-shielding effect of lumbar interbody fusion devices”, Journal of Neurosurgery 93:259-65, 2000), leading to subsequent decreased mineralization, bone resorption, and significant bone mineral density decrease in long-term (see Cunningham et al., “A quantitative densitometric study investigating the stress-shielding effects of interbody spinal fusion devices: Emphasis on long-term fusions in thoroughbred racehorses”, Trans Orthop Res Soc 23:250, 1998).
Many current efforts to reduce these complications have concentrated on using poly (α-hydroxy) ester polymers that have much lower stiffness than metallic materials to fabricate conventional cage designs (see, Kandziora et al., “Biomechanical analysis of biodegradable interbody fusion cages augmented with poly(propylene glycol-co-fumaric acid)”, Spine 27:1644-51, 2002; Toth et al., “Evaluation of 70/30 poly (L-lactide-co-D,L-lactide) for use as a resorbable interbody fusion cage”, Journal of Neurosurgery 97:423-32, 2002; van Dijk et al., “Bioabsorbable poly-L-lactic acid cages for lumbar interbody fusion: three-year follow-up radiographic, histologic, and histomorphometric analysis in goats”, Spine 27:2706-14, 2002). Degradable cages possess a number of significant advantages over non-degradable materials including eventual removal of all foreign material that could cause nerve root irritation, alleviation of stress-shielding effects and reduce adjacent level disc disease, and removal of imaging artifact. Nevertheless, the mere replacement of base material from original designs might lead to cages that cannot provide adequate stability since biodegradable polymers have less stiffness/strength than permanent materials and this reduced stiffness/strength will be further compromised over the degradation time. Furthermore, primary degradation products of these poly (α-hydroxy) acids form a low pH environment that can inhibit osteogenesis. It has been shown that even small pH shifts can significantly affect bone marrow stromal cell (BMSC) function of proliferation and differentiation (see Kohn et al., “Effects of pH on human bone marrow stromal cells in vitro: Implications for tissue engineering of bone”, Journal of Biomedical Materials Research 60:292-9, 2002) since the growth and development of osteoblasts are linked to regulation of pH and acidity of the extracellular microenvironment (see Chakkalakal et al., “Mineralization and pH relationships in healing skeletal defects grafted with demineralized bone matrix” Journal of Biomedical Materials Research 28:1439-43, 1994; Green “Cytosolic pH regulation in osteoblasts”, Mineral and Electrolyte Metabolism 20:16-30, 1994; Kaysinger et al., “Extracellular pH modulates the activity of cultured human osteoblasts”, Journal of Cellular Biochemistry 68:83-913-15, 1998). Therefore, although degradable polymer cages offer significant potential advantages over non-degradable cages, there are also significant hurdles to overcome including the maintenance of adequate mechanical properties and reduction of acidic degradation products.
With various bone graft substitutes emerging as biological inducers to achieve successful arthrodesis, delivery within a restricted volume becomes critical. Among a variety of promising bone graft substitutes, bone growth factors and cell-based approaches particularly require suitable delivering vehicles (see Helm et al. “Bone graft substitutes for the promotion of spinal arthrodesis”, Neurosurg Focus 10:1-5, 2000). Several recombinant human bone morphogenic proteins (rh-BMPs) have been approved for certain clinical applications, and they are commonly delivered through an absorbable collagen sponge to effectively achieve arthrodesis by osteoinduction. However, current delivering approaches are associated with the inability to directly deliver bone morphogenic proteins for bone regeneration.
The only commercially available delivery system at this moment for bone morphogenic protein consists of collagen sponges soaked in bone morphogenic protein solutions that contain bone morphogenic protein at concentrations over a million times higher than what is physiologically found in the human body. The release of the bone morphogenic protein in this fashion obviously consists of a very large bolus quantity of which all its effects are unknown. Reports have shown that bone morphogenic protein can also cause an initial osteolysis of surrounding bone secondary to what is thought to be an initial drain of osteogenic cells from surrounding bone towards the bone morphogenic protein soaked sponges. This can initially weaken surrounding bone structures thus promoting subsidence of any supporting implants. Furthermore, a high concentration of bone morphogenic protein has been shown to cause swelling of surrounding soft tissue with resultant swallowing and breathing difficulty. Another disadvantage of uncontrolled release of bone morphogenic protein is the ectopic formation of bone. Bone formation distal from the intended site of osteogenesis can result in radiculopathy as well as intradural bone formation.
Primary requirements in developing biodegradable cages are assuring that porous degradable cages can withstand surgical impaction forces, can carry in vivo spinal forces initially and up to the time bony fusion is achieved (normally 3-6 months), and have degradation products that will not adversely affect bone regeneration. However, bone tissue engineering within degradable constructs invokes two new requirements in addition to the primary degradable cage requirements delineated above. The first is osteoconductivity, which is the ability to promote and support ingrowth of bone-forming cells. Among the most common strategies to confer osteoconductivity to an orthopedic implant material involves coating with a calcium-phosphate-based mineral film similar to bone mineral. These films have a well-characterized positive effect on the ingrowth and proper function of bone-forming cell types, including osteoblasts and osteoblast precursors. U.S. Pat. No. 6,767,928 (which is incorporated herein by reference as if fully set forth herein) shows that calcium-phosphate mineral coatings can be grown on porous polymer scaffolds, and that the mineral coatings positively influence bone tissue growth. The technology used to grow these mineral coatings mimics the process of natural bone mineralization, and the coatings have a structure, mineral phase, and elemental composition that is similar to human bone mineral (see also, Bunker et al., “Ceramic thin film formation on functionalized interfaces through biomimetic processing”, Science 264:48-55, 1994; Mann et al., “Crystallization and inorganic-organic interfaces—biominerals and biomimetic synthesis”, Science 261:1286-92, 1993; Murphy et al., “Bioinspired growth of crystalline carbonate apatite on biodegradable polymer substrata”, J Am Chem Soc 124:1910-7, 2002; and Ohgushi et al., “Stem cell technology and bioceramics: from cell to gene engineering”, J Biomed Mater Res 48:913-27, 1999). These “bone-like” mineral coatings have been shown to significantly enhance osteoconductivity of orthopedic implant materials (see Ohgushi et al.; Hench, “Bioceramics: From concept to clinic”, Journal of the American Ceramic Society 74:1487-510, 1991; Murphy et al., “Bone regeneration via a mineral substrate and induced angiogenesis”, J Dent Res 83:204-10, 2004). In addition to their osteoconductivity, mineral coatings also represent a potential vehicle for delivery of osteogenic growth factors (see Seeherman et al., “Bone morphogenetic protein delivery systems”, Spine 27:S16-23, 2002). Multiple bone growth factors, including BMP-2, IGF-1 and TGF-β (see Gittens et al., “Imparting bone mineral affinity to osteogenic proteins through heparin-bisphosphonate conjugates”, J Control Release 98:255-68, 2004; Gorski et al., “Is all bone the same? Distinctive distributions and properties of non-collagenous matrix proteins in lamellar vs. woven bone imply the existence of different underlying osteogenic mechanisms”, Crit Rev Oral Biol Med 9:201-23, 1998; Gorski et al., “Bone acidic glycoprotein-75 is a major synthetic product of osteoblastic cells and localized as 75- and/or 50-kDa forms in mineralized phases of bone and growth plate and in serum”, J Biol Chem 265:14956-63, 1990; Liu et al., “Bone morphogenetic protein 2 incorporated into biomimetic coatings retains its biological activity”, Tissue Eng 10:101-8, 2004; Matsumoto et al., “Hydroxyapatite particles as a controlled release carrier of protein”, Biomaterials 25:3807-12, 2004; and Sachse et al., “Osteointegration of hydroxyapatite-titanium implants coated with nonglycosylated recombinant human bone morphogenetic protein-2 (BMP-2) in aged sheep”, Bone 37:699-710, 2005) have been shown to interact strongly with bone-like mineral substrates. Therefore, it is possible that calcium phosphate mineral substrates can be coated with growth factors, and these factors can subsequently be presented to bone-forming cells growing into a scaffold construct. Previous studies have demonstrated that it is indeed possible to use hydroxyapatite minerals as template substrates to bind and release bone growth factors, particularly BMP-2, and that the bound growth factors induce bone ingrowth in vivo (see Gittens et al.; and Sachse et al.).
Notwithstanding the foregoing advances in tissue engineering, there is still a need for improved cages for facilitating the fusion of adjacent bones such as vertebrae, or adjacent bone surfaces such as in an open fracture.